Over 12 million new cancer cases are diagnosed worldwide each year. About 40% of all cancer patients receive radiation therapy as part of their curative treatment. Most types of radiotherapy use photons (X-rays or gamma-rays) or electron beams for the local treatment of disease. Ionizing radiation damages the DNA of tumour and healthy cells alike, triggering complex biochemical reactions that eventually result in prolonged abnormal cell function and cellular death.
The aim of radiotherapy treatment is to maximize the absorbed dose (and hence damage) to the target tumour and to minimize radiation-induced morbidity to adjacent healthy tissue. This is generally achieved by targeting a beam of radiation at the tumour area along a path that spares nearby critical and radiosensitive anatomic structures. In some known arrangements, multiple beams may be employed each travelling along a different path, the paths being arranged to cross one another in the tumour region. This has the advantage of avoiding overexposing the same healthy tissues. The total radiation dose to be delivered to a tumour may be partitioned into fractions over successive sessions. Because healthy tissues recover better and faster than malignant ones, with each radiotherapy session the accumulated cellular damage in the targeted tumour increases, whilst normal (non-tumour) tissues are given the opportunity to repair.
The absorbed dose of radiation as a function of depth in human tissue is illustrated in FIG. 1 for X-ray radiation of energy 4 MeV (trace T1), X-ray radiation of energy 20 MeV (trace T2) and proton radiation of energy 150 MeV (trace T3). It can be seen clearly from FIG. 1 that proton radiation exhibits the sharpest peak in energy deposited as a function of depth, and that a beam of energy 150 MeV is able to deposit a substantial amount of energy at a depth in the range corresponding to that of the tumour site, the depth being between depth d1 (around 10 cm) and depth d2 (around 12 cm) in the example of FIG. 1, with relatively low amounts of energy deposited at depths outside of this range.
When irradiating beams are composed of heavy charged particles (protons and other ions, such as carbon), radiation therapy is generally called hadrontherapy. If protons are used, radiation therapy may be called proton therapy. The strength of hadron therapy lies in the unique radiobiological properties of these particles. The particles can penetrate tissue, and hadrons deposit their maximum energy just before stopping within the tissue. This allows a precise definition of the specific region to be irradiated. The peaked shape of the proton (or hadron) energy deposition as a function of distance in the tissue is called the Bragg peak as indicated at BP in FIG. 1 and FIG. 2. FIG. 2 is a plot of deposited dose as a function of depth for proton beams of energy 50 MeV (trace T1), 150 MeV (trace T2) and 200 MeV (trace T3). With the use of protons and other hadrons, a tumour at a given depth can be subject to a substantial deposited dose of energy by means of protons or other hadrons whilst the damage to healthy tissues as a consequence of irradiation of the tissue is less than in the case when x-rays are employed as illustrated in FIG. 1.
The depth of tissue traversed prior to the depth at which the Bragg peak is found is set by the energy of the incident proton beam as illustrated in FIG. 2. The width of this peak can be moderated by manipulating the range of energies of the protons or other charged hadrons that comprise the irradiating beam.
It is to be understood that the Bragg peak occurs immediately before charged hadron particles come to rest. The peak in energy loss by a charged particle as it moves through a material occurs because the interaction cross section increases as the energy of the particle decreases.
It is to be understood that the position of the Bragg peak can be adjusted by means of an attenuator which absorbs a portion of the energy of a particle, or by modifying the properties of the particle accelerator. By varying the amount of attenuation in real time, the Bragg peak associated with an otherwise monoenergetic proton beam (exhibiting a relatively sharp Bragg peak) may be effectively widened over a given time period by increasing the range of energies, so that a larger volume of tissue (for example tumour tissue) can be treated. Real-time adjustment of the amount of attenuation can be achieved by movement of a variable thickness attenuator such as by rotation of a wedge-shaped attenuator forming part of a spinning wedge attenuator device.
As a consequence of the relatively sharp Bragg peak immediately prior to charged hadrons coming to rest, tissues closer to the surface of the body than the tumour site receive much reduced radiation when protons or other hadrons are employed, and therefore reduced damage. Tissues deeper than the tumour within the body receive very few hadrons, so that the hadron dose becomes immeasurably small.
The advantages of hadron therapy, such as proton therapy, through its ability to deliver very high doses into tumours with much reduced dose to neighbouring tissues, may include at least one of reduced probability of second cancers, increased ability to treat tumours adjacent to critical organs or structures, a reduction in overall treatment time, and improved quality of life for patients during and after treatment.
FIG. 3 is a schematic illustration of a known proton beam computerised or computed tomography (CT) scanner 100. The scanner 100 has first and second beam tracker structures 110, 120 and a calorimeter device 170, each of which is in communication with a computing device 190. The scanner 100 is arranged to allow a subject to be positioned between the first and second beam tracker structures 110, 120. A beam of protons 101B is projected towards the subject from a source (not shown) through the first beam tracker structure 110 to the subject 101S. Protons emerging from the subject 101S pass through the second beam tracker structure and into the calorimeter device 170.
The beam tracker structures 110, 120 each have a pair of mutually parallel beam position-sensitive detectors 110A, 110B, 120A, 120B configured to detect a location within a 2D X-Y plane defined by each detector 110A, 110B, 120A, 120B at which the beam 101B passes through the detector 110A, 110B, 120A, 120B. By knowing the position within the 2D planes defined by each of the detectors 110A, 110B of the first beam tracker structure 110 at which the beam 101B passes through the computing device 190 of the apparatus 100 is able to calculate a vector v1 defining the path of travel of the beam 101B from the source to the subject 101S.
Similarly, knowing the position within the 2D planes defined by the detectors 120A, 120B of the second beam tracker structure 120 at which the beam 101B passes through the detectors 120A, 120B the computing device 190 is able to calculate a vector v2 defining the path of travel of the beam 101B from the subject 101S to the calorimeter device 170.
The calorimeter device 170 is configured to measure the amount of energy contained in the beam 101B entering the device 170. The device 170 provides an output signal to the computing device 190 indicative of the amount of energy contained in the beam 101B at a given moment in time. In the example shown the calorimeter is a CsI-based scintillator calorimeter device 170.
The computing device 190 is configured to correlate measurement of vectors v1 and v2 with a measurement of the energy of the beam 101B as determined by the calorimeter device 170. It is to be understood that, based on a knowledge of the energy of protons incident on the subject 101S, the vectors v1 and v2 and the energy of the beam 101B emerging from the subject 101S, the computing device 190 is able to calculate an amount of proton energy absorbed by the subject 101S (i.e. the dose) at a given location within the subject 101S in a known manner.
In the scanner 100 shown in FIG. 3, the subject 101S is rotated about an axis A through the subject 101S parallel to the Y-axis as shown in FIG. 3. Proton intensity data is captured as a function of rotational position of the subject 101S about the A-axis, and the computing device 190 is able to build up a 3D image of the fraction of proton energy absorbed at a given 3D location within the subject 101S. The scanner 100 of FIG. 3 is a ‘broad beam’ scanner 100 in that the beam 101B is arranged to irradiate substantially continuously the area being imaged, in contrast to scanned beam systems in which the beam 101B is scanned in the X-Y plane.
Since different tissues exhibit different absorption characteristics, the internal structure of the subject's anatomy can be determined from the 2D images (radiographs) and 3D datasets built up from the 2D images captured as a function of rotational position of the subject 101S about axis A.
It is to be understood that, knowing the incident energy of a proton, and tracking it through the apparatus so as to determine its residual energy following passage through the tissue, allows an absorbed dose of proton radiation to be calculated. Additionally, tracking the paths of individual protons over a range of incident angles allows the reconstruction of the 3-dimensional volumetric CT image as described in further detail below. It is to be understood that measuring the energy of each proton and tracking the path of the proton so as to calculate where in the subject the proton lost its energy is important in some embodiments. This is because charged hadrons such as protons are typically relatively strongly scattered by the subject compared with X-rays. In contrast, in the case of X-ray CT scanner systems it is not necessary to measure the exit energy of each X-ray in order to generate a CT image of a subject.
The first proton CT experiment and reconstruction was performed in 1976 (A. M. Cormack and A. M. Koehler, “Quantitative Proton Tomography: Preliminary Experiments,” Phys. Med. Biol. 21, 560-569, 1976). Little further development was made until about 2000, when the challenge of producing clinically viable proton CT was taken up. The Paul Scherrer Institute (Switzerland) published details of their system (P. Pemler, J. Besserer, J. de Boer, M. Dellert, C. Gahn, M. Moosburger, U. Schneider, E. Pedroni, H. Stauble, “A detector system for proton radiography on the gantry of the Paul-Scherrer-Institute,” Nuclear Instruments and Methods in Physics Research A432:483-495, 1999)), which used single-layer segmented scintillating fibers and a range telescope constructed from closely-packed plastic scintillator plates. With only one plane of detectors before and after the object being investigated, they could only detect the position of an individual proton and not its direction. The use of scintillating plates in a range telescope for measuring proton beam energy post-passage through tissue meant that only one proton could be unambiguously detected at one time.
Loma Linda University, USA, developed a proton CT system having four x-y resolving silicon strip detectors, two positioned before the patient and two after, in a similar manner to that shown in FIG. 3. The range telescope was a Cesium Iodide-based scintillator calorimeter comprising 18 crystals with the resultant light detected by a photodiode connected to each crystal (H. F. W. Sadrozinski, V. Bashkirov, B. Colby, G. Coutrakon, B. Erdelyi, D. Fusi, F. Hurley, R. P. Johnson, S. Kashiguine, S. McAllister, F. Martinez-McKinney, J. Missaghian, M. Scaringella, S. Penfold, V. Rykalin, R. Schulte, K. Schubert, D. Steinberg, A. Zatserklaniy, “Detector Development for Proton Computed Tomography (pCT)”, in Conference Proceedings of IEEE Nuclear Science Symposium and Medical Imaging Conference, 2011). The system was very slow, taking several hours to obtain one scan. The system also had limited energy resolution of the range telescope due to the use of a calorimeter and the thickness of strip detectors which perturb the proton beam. More recently, the same group announced the development of a second system, again using silicon strip detector pairs either side of the patient but with a range telescope having a stack of polystyrene scintillators (3 to 10) read out by photomultipliers (H. F.-W. Sadrozinski, R. P. Johnson, S. Macafee, A. Plumb, D. Steinberg, A. Zatserklyaniy, V. A. Bashkirov, R. F. Hurley, R. W. Schulte, “Development of a head scanner for proton CT,” Nuclear Instruments and Methods in Physics Research A699:205-210, 2013). These systems suffer from limited energy resolution due to the use of a low number of scintillating planes in the range telescope and the use of relatively thick silicon strip detectors, which adversely affect the quality of the incident proton beams entering the patient. Moreover, the present applicant has recognized that the limitations of two conventional crossed x-y strip detectors will create a high proportion of false events. This is because with N events (protons) detected within one read cycle, there will be N(N−1) false events recorded.
A further system has been developed that apparently overcomes some of the range telescope limitations in terms of energy resolution by using 48 thin plastic scintillators coupled to silicon photon multipliers (M. Bucciantonio, U. Amaldi, R. Kieffer, F. Sauli, D. Watts, “Development of a Fast Proton Range Radiography System for Quality Assurance in Hadron therapy,” Nuclear Instruments and Methods in Physics Research A: http://dx.doi.org/10.1016/j.nima.2013.05.110). As currently described the system is unsuitable for proton CT as no provision is made for positional detectors prior to the patient being imaged.
US2013/0015352A1 describes a proton computed tomography (pCT) detector system, including two tracking detectors in sequence on a first side of an object to be imaged, two tracking detectors in sequence on an opposite side of the object to be imaged, a calorimeter, and a computer cluster, where the tracking detectors include plastic scintillation fibers. All fibers in the detector system are read out by silicon photomultipliers.
A limitation of range telescopes which comprise of sheets of scintillator is that they are unable to distinguish between multiple protons passing through the sheets within the resolving time of the system, and there is no information provided about the location of the proton within the area of the sheet. Reducing the flux of protons to such a low level so that normally there is only one proton passing through the range telescope at one time, in order to compensate for this deficiency, means that the time to record a satisfactory CT image can be excessively long.
It is desirable to provide improved apparatus for delivery of radiotherapy treatment. It is an aim of the present invention to address disadvantages associated with the prior art.